Femoral stems with reduced stiffness have been introduced in total hip arthroplasty to facilitate proximal load transfer and thereby reduce stress shielding and periprosthetic bone loss. However, poor implant fixation and unacceptably high revision rates are a major problem with these prostheses. One reason for this is that the implant is precisely machined and the femoral canal is frequently not, leaving gaps as large as 0.025″ between the implant and the wall of the femoral canal. In many instances the implants may only have 35% of their surface area in direct contact with the adjacent bone. See FIG. 1. This lack of a tight fit between the implant and the surrounding bone is a significant problem, inasmuch as a tight fit is required between the implant and the adjacent bone in order to provide maximum fixation in the shortest time, by maximizing implant stability and the opportunity for bone ingrowth.
For successful implants, sufficiently regenerated bone fills the gap between the implant and the host bone, so that the implant is firmly attached to the surrounding bone.
To overcome problems with implant loosening, implants need to stimulate rapid bone regeneration in order to replenish the missing bone and/or to fix the implant firmly within the host bone. To succeed as an orthopedic implant, the implant must provide a habitat for bone-forming cells (e.g., osteoblasts) so that the bone-forming cells can colonize on the implant surface and synthesize new bone tissue. Frequently the implants are not compatible with the bone cells responsible for bone formation, and instead promote the formation of undesirable fibrous soft tissue. Such fibrous soft tissue does not adequately support the implant, which leads to implant loosening under physiological loading conditions and eventual implant failure. Thus, in order to design more successful orthopedic implants, one needs to take into account the cellular processes that promote bone ingrowth. Positive responses from osteoblasts, including increased initial adhesion, proliferation and differentiation (from noncalcium-depositing cells to calcium-depositing cells) are essential. Coordinating activities between osteoblasts and the bone-resorbing cells (e.g., osteoclasts) is also needed in order to provide healthy bone around the implant. Poor communication between cells can lead to bone necrosis adjacent to the implant, thereby causing loosening of the implant. Another undesirable occurrence is the formation of fibrous soft tissue by fibroblasts. Excessive fibrous soft tissue formation hinders osteoblast/osteoclast activities and hence limits bone regeneration. Due to these cellular events, the orthopedic field has concentrated on understanding cellular recognition of surfaces and creating biomaterial surface properties which maximize such interactions for the creation of more bone and enhanced osseointegration.
One way to improve the performance of bone implants is to modify the surface texture of the implants. Many studies have shown that microstructural features such as grain and particle size promote osteoblast functions better than smooth surfaces. This motivates the use of nanophase materials for orthopedic implants.
Macrostructural features such as porous coatings are another means for improving osseointegration of the implant. Today, hip implant stems are typically a composite structure consisting of a substrate (typically formed out of a cobalt chrome alloy or a titanium alloy) which carries the patient's weight, and a porous surface coating mounted on the implant substrate. This porous surface coating (which is generally referred to in the industry as a “porous coating”) comprises peaks and valleys, whereby to aid in immediate implant fixation and ultimately promote long term stability through osseointegration of the host bone with the porous coating. See FIG. 2.
Prior to inserting the implant, the surgeon broaches the femoral canal to create a cavity that, ideally, closely matches the geometry of the implant (which is then inserted into the cavity in the bone). However, this fit is not always perfect, and gaps frequently exist between the implant and the bone. These gaps cause the implant to “point load” the surrounding bone, and also create barriers which inhibit rapid and effective osseointegration of the implant.
Today, the majority of porous coatings are made of titanium or tantalum. These porous coatings are “static”, in the sense that they are substantially rigid. These porous coatings are textured, and are applied to the implant substrate by hot plasma spray, chemical and/or physical vapor deposition, by chemically etching thin films and plates, and/or by sintering and/or diffusion bonding metal beads or metal fibers into a solid rigid mass. See FIG. 3.
In addition to producing a substantially rigid structure, the coating processes used to produce porous coatings tend to produce a largely two-dimensional structure for the bone to grow around. There is no means for the bone to tunnel further into the porous coating so as to establish significant three-dimensional osseointegration. Thus, the largely two-dimensional porous coating may stifle or compromise effective long-term osseointegration of the implant due to the lack of significant three-dimensional osseointegration. Additionally, the largely two-dimensional porous coating structures created using these prior art technologies do not accurately mimic the structure of trabecular (i.e., cancellous) bone, which is three-dimensional and includes interconnecting networks of pores with capillarity properties. See FIG. 4.
Recently, there have been advances in the creation of porous coatings that more accurately resemble trabecular bone. These porous coatings have interconnecting networks of pores which are similar to those of trabecular bone, and may serve to promote bone ingrowth deeper into the porous coating and hence provide better long-term implant fixation. One method known in the art for creating such a porous coating is through the replication of an open cell network. In this method, a structure similar to trabecular bone (e.g., a polyurethane foam) is coated with another material (e.g., titanium or tantalum) by vapor deposition, low temperature arc vapor deposition (LTAVD), chemical vapor deposition, ion beam assisted deposition and/or sputtering. The underlying structure (e.g., the polyurethane foam) may then undergo pyrolysis so as to remove the underlying structure (e.g., the polyurethane foam), leaving a metallic structure which can be attached to the hip implant substrate (e.g., by sintering, brazing, diffusion bonding, gluing or cementing, etc.). See FIG. 5. However, the porous coating produced by this method is static, i.e., it is substantially rigid.
Other methods for forming a porous coating include chemical vapor deposition of commercially pure tantalum onto a porous carbon scaffold and then sintering the resulting structure onto the substrate of the implant. See FIGS. 6 and 7. Again, however, the porous coating produced by this method is static, i.e., it is substantially rigid.
Depending on the desired thickness of the struts of the porous coating, the physical or chemical vapor deposition process may be sufficient to reproduce the scaffold structure; however, it is also possible to more rapidly thicken the struts following deposition through the bulk application of a powdered metal or granulates made of titanium, tantalum or other biomaterials, with or without a binder (e.g., methylcellulose). Following the application of the powdered metal, the scaffold is sintered to integrate the powder or granulates. See FIG. 8. In any case, however, the porous coating produced by this method is static, i.e., it is substantially rigid.